Device for swept-source optical coherence domain reflectometry

ABSTRACT

A device for swept-source optical coherence domain reflectometry (SS OCDR) on moveable samples, particularly human eyes, for obtaining A-scans, having a measuring range according to the sample length and having a laser light source which can be adjusted by a main wave number k 0  and at least one receiver for the light dissipated from the sample, wherein the sample is illuminated on the sample surface by a measurement beam having a diameter D by way of a coupling device. The light source has a spectral line width of δk&lt;168 m −1  and the adjustment of the light source is carried out in τ&lt;44 s/(D*k 0 ).

PRIORITY CLAIM

The present application is a National Phase entry of PCT Application No.PCT/EP2009/009189, filed Dec. 21, 2009, which claims priority fromGerman Application No 102008063225.2, filed Dec. 23, 2008, thedisclosures of which are hereby incorporated by reference herein intheir entirety.

FIELD OF THE INVENTION

The invention relates to a device for swept source optical coherencedomain reflectometry such as can be applied in the optical biometry ofthe eye, for example.

BACKGROUND

The optimal matching of an artificial intraocular lens requiresknowledge of the optical conditions in the patient's eye, in particularthe distances between the cornea, crystalline lens and retina.

After this determination of position was originally carried out by meansof ultrasound, a device operating optically and without making contacthas been introduced in the form of the IOL Master of Carl Zeiss. Thefunctional principle is based in this case on the so called time domainoptical coherence domain reflectometry, a short coherence interferometrymethod such as is described, for example, in WO 00/33729, the content ofwhich is incorporated by reference herein. The main component is aMichelson interferometer that enables the detection of interference oflight scattered back by the cornea, lens and retina. The use of a shortcoherence light source means that it is possible for always only shortwave chains to interfere with one another, and this determines themeasuring accuracy. So that axial movements of the patient do notfalsify the measurement result, the so called dual beam method isapplied in which the light scattered back by the cornea serves asreference.

Since the measuring range that must be more than 43 mm for an eye(typical eye lengths vary between approximately 20 and 32 mm, extremeones between 14 and 40 mm, the mean refractive index being approximately1.36), must be traversed mechanically by the reference mirror in thecase of Michelson interferometer, a measurement usually lasts a fewseconds in which the patient is, for example, not allowed to blink sincethe eyelid movement would render the measurement impossible.

Efforts to accelerate the rate of adjustment of the reference path, forexample, by rotating prisms such as EP 1 391 781, have not beensuccessful, since the sensitivity is not sufficient to achieve therequired measuring accuracy.

In DE 43 09 056 describes another measurement method based on shortcoherence, in the case of which light from a broadband light source isshone into the sample, and the light scattered back from various depthsis analyzed spectrally. The depth information is obtained from a Fouriertransformation of the detected signal. This method is denoted asspectral domain OCDR (SD ODCR) or, because of the Fourier transformationused, also as Fourier domain OCDR (FD OCDR). This category also includesthe swept source OCDR (SS OCDR), which is described in the articleentitled “High-speed optical frequency-domain imaging” by S. H. Yun etal., Optics Express 2003, page 2953, and in which the light source istuned spectrally, and the signal received by the detector likewiseincludes the depth information after the Fourier transformation. Asalready shown in U.S. Pat. No. 5,321,501 for time domain OCT (TD OCT),the imaging required to implement optical coherence tomography (OCT) isimplemented by Galvo scanners that deflect the measurement beamlaterally over the sample.

Along the lines of the terminology introduced in the case of theultrasound measuring device, the one-dimensional (axial) measurement inthe case of OCDR along the light axis is generally denoted as an A-scanin general, and therefore also below. Likewise along the lines of theultrasound terminology, the two-dimensional measurement with the aid ofa lateral component in the case of OCT is also denoted as a B-scan.

A first attempt to apply SS OCDR in optical biometry was described in F.Lexer, C. K. Hitzenberger, A. F. Fercher and M. Kulhavy“Wavelength-tuning interferometry of intraocular distances”, Appl.Optics 36 (1997) pages 6548-6553. This solution showed that it ispossible in principle to measure the intraocular distances in the eye,although the measuring accuracy was much too inaccurate at 0.82 mm.

An improvement to this solution was disclosed in C. K. Hitzenberger, M.Kulhavy, F. Lexer, A. Baumgartner “In-vivo intraocular ranging bywavelength tuning interferometry”, SPIE [3251-6] 1998. Here, aresolution of 0.15 mm was reached, but it still does not correspond tothe requirements. The measuring accuracy for the eye length must,however, be smaller than 30 μm in order to limit the residual errors ofthe determined IOL refraction to 1/10 diopters.

In particular, the OCDR and OCT methods on moving samples such as, forexample, the human eye have the problem that the sample can move duringthe measurement and this, as discussed in S. H. Yun et al. (2004),OPTICS EXPRESS 2977, can greatly reduce the signals and falsify them.The usual approaches to eliminating the problem are the extremelycomplicated tracking methods in which the movement of the sample isdetected and the measurement beam position is tracked.

Such approaches to the compensation of typical movements of a fewhundred micrometers per second are described, for example, in Hammer etal. (2005), Journal of Biomedical Optics 10(2), 024038, and in US2006/105903. It is disadvantageous of such approaches that, despite thelarge technical outlay, the finite latency time of such systems alwaysresults in certain tracking errors, particularly for very fast eyemovements such as saccades.

SUMMARY OF THE INVENTION

Proceeding from the prior art, it is therefore the object of theinvention to specify a device with the aid of which the intraoculardistances in the eye can be measured quickly and accurately,particularly even given the occurrence of typical eye movements, withoutexhibiting the disadvantages of an active measurement beam tracking witha sample movement such as, for example, latency time errors.

This object is achieved by a device for swept source optical coherencedomain reflectometry (SS OCDR) on movable samples, particularly humaneyes, for the purpose of obtaining A-scans, having a measuring rangecorresponding to the sample length and having a laser light sourcetunable about a centroid wave number k₀ and at least one receiver forthe light backscattered from the sample, the sample being illuminatedvia a coupling device on the sample surface with the aid of ameasurement beam of diameter D, in that the light source has a spectralline width of δk<168 m⁻¹, and in that the tuning of the light source isperformed in τ<44 s (D*k₀).

Distances are therefore measured with low expenditure and efficientlyover the entire length of the eye since, despite typical eye movementsof up to 1000 μm/s and given only moderate requirements for the tuningrate of the source, determine interfering signal losses resulting fromsample displacements are avoided in the case of distance measurementsbetween surfaces of the crystalline lens, cornea and retina.

The result of the inventive solution is therefore that the tuning timeof the laser is matched to the sensible measurement beam profiles in thesample so that the lateral sample displacements possible during thetuning time of the laser can predominantly amount only to fractions ofthe smallest possible measurement beam diameter in the sample.Interfering signal losses, in turn, are therefore avoided by averagingout different lateral interference modulations, since the sample volumesilluminated at various instances during the tuning of the measurementbeam have a sufficient overlap.

What is understood here by sensible measurement beam profiles are thosethat can supply signal strengths of cornea, crystalline lens and retinathat suffice for spacing apart by virtue of the fact that they havemeasurement beam waist positions in the region between the back of thecrystalline lens of relatively short eyes as far as to the retina oflong eyes (8 . . . 40 mm).

At the same time, the inventive solution also avoids a signal reductionand signal corruption through axial sample displacements during tuning,for example extension or compression of the A-scan with resultingunacceptable errors in the spacing apart of the eye structures.

Thus, it is predominantly undisturbed signals that result, without theneed for active tracking of the measurement beam with sample movements.

It is advantageous in this case when the light source has a spectraltuning range Δk about a centroid wave number k₀ of at least Δk>18 000m⁻¹.

In this case the ratio of the tuning range Δk and line width δk isadvantageously greater than 360, further preferably greater than 2000,further preferably greater than 4000, and yet further preferably greaterthan 9000. This ratio ensures the implementation of an adequate ratiobetween the measurement depth and measurement resolution.

A further advantage results when the quotient of the tuning rate (Δk/τ)and laser line width δk is greater than 18 kHz, preferably also greaterthan 4 MHz, with further preference greater than 40 MHz.

It is particularly advantageous in this case when the detection of thelight backscattered by the cornea, crystalline lens and retina isperformed during a single tuning of the light source. It is advantageoushere when the backscattered light detected at the receiver is digitizedat a rate of more than Δk (τ*δk). This ensures that the spectralinformation is adequately scanned.

The inventive device is particularly suitable when the line width δk ofthe light source lies between 22 and 50 m⁻¹. Such line widths can beimplemented, for example, with tunable fiber ring lasers, and offeracceptable drops in sensitivity over the measurement depth.

It is advantageous when the bandwidth of the at least one receiver,described, for example, by the cutoff frequency with 3 dB signal drop,is greater than 2*Δk/(τ*δk) and preferably less than 80 MHz.

According to the invention, the device for SS OCDR on the eye is atunable laser source, an interferometer with a reference arm and asample arm closed off from the sample, detectors at the interferometeroutputs, and a signal processing unit for the detected signals.

According to the invention, the position of retina signals and corneasignals in the A-scan, and the laser line width δk are matched to oneanother in the device for SS-OCDR on the eye.

It is also particularly preferred to use a reference interferometer forwave number references of the source, and thus for length calibration ofthe OCDR signal.

It is particularly advantageous for the device for SS OCDR on the eyewhen the measurement beam diameter D is smaller than 3 mm in the regionof the sample entrance.

It is also preferred for the measurement beam to be convergent beforeentering the eye, it preferably being possible to set or switch over thesize of the convergence. Particularly, by setting the convergence it ispossible to match to one another, or to optimize the relative strengthof the detected signal from various eye regions (retina, lens, cornea)to facilitate a joint processing for measuring distance and visualevaluation of the signals.

Alternatively, it can also be advantageous when the measurement beam iscolliminated before entering the eye, and means are provided forrefixation the eye, in order to be able to detect specula cornea signalsand lens signals. Owing to the refixation, it is possible to measure onthe optical axis of the eye, which can deviate by up to 14° from thevisual axis.

It is also advantageous when it is possible to switch over between acolliminated and a convergent measurement beam.

A further advantageous implementation of the invention results when themeasurement beam strikes the eye outside the corneal apex, it beingpreferred to provide for the purpose of positioning the measurement beamrelative to the eye an apparatus that can be driven, in particular, byevaluating the light detected by the receiver. This prevents a strongreflection by the corneal apex from saturating the detectionapparatuses, or else prevents the signal-to-noise ratio from beingdegraded, for example by the increase in the shot noise component causedby the strong cornea reflection.

Such an apparatus for positioning the measurement beam, particularlywhen synchronized with the tuning of the laser, permits B-scans to beobtained and therefore permits the implementation of OCT over the entireeye. Such positionings of the measurement beam can, for example, beperformed by optical configurations known per se in combination withangle deviations by galvanometer scanners, which are likewise known(U.S. Pat. No. 5,321,501).

It is particularly advantageous when a photon flux of 3*10⁸ to 1*10¹³ isdirected onto the sample in the tuning time τ of the light source.Phototoxic effects can be avoided by suitable selection of thewavelength (for example 1016 nm). It is thereby possible to implement ahigh sensitivity for the weakly scattering eye structures by optimizingthe signal-to-noise ratio caused by the shot noise, without there beingdamage to the eye tissue.

It can, furthermore, be advantageous when the measurement is performedalong the visual axis of the eye since, in particular, the distanceinformation on this axis is very valuable for matching intraocularlenses. To this end, the measurement beam and fixer light should strikethe eye collinearly, but they can also have different divergences, forexample in order to compensate a possible refraction error of the eye.

An advantageous refinement of the invention results when monomode fibersare used in the reference arm and/or source arm and/or detection armand/or reference interferometer, in order to avoid disturbing parasiticinterferences between different modes propagating in the fiber andartifacts resulting therefrom. Equally, the open fiber ends are designedas obliquely polished surfaces in order to avoid interferingretroreflections.

It is advantageous when the reference signals and sample signals aredigitized at a constant scanning rate, it being preferred to use thesame scanning rates for reference signals and sample signals. However,in order to reduce the data volume that occurs, it is also possible,given a suitable selection of the reference interferometer, to scan thereference signal at a lower scanning rate than the scanning of thesample signal.

According to the invention, the measurement beam has a wavelength ofbetween 600 and 1150 nm, wavelengths of 700 nm, 800 nm and 1060 nm beingparticularly preferred.

A preferred refinement of the device for SS OCDR on the eye results inthe fact that an apparatus is provided for projecting onto the eye,particularly the cornea, aiming markers with a wavelength of between 400and 1500 nm and an observation unit is provided for detecting thereflections of these aiming markers. These can also be evaluated withregard to the determination of the position and shape of the cornea andlens.

It is further advantageous in this case when the observation unit, forexample a camera, is provided for checking the adjustment of themeasurement beam relative to the eye, the observation unit preferablybeing sensitive to the wavelengths of the measurement beam and theaiming markers. Cameras with a silicon sensor seem to be particularlysuitable here because of their adequate residual sensitivity in the nearinfrared.

A particularly suitable refinement of the invention results when thelight source is movable relative to the eye, light source and referenceinterferometer preferably being firmly connected.

Another advantageous condition for implementing the invention consistsin that the interferometer is movable relative to the eye, light sourceand interferometer preferably being firmly connected.

It is advantageous in this case when the electrical and opticalconnections are designed to be separable.

The inventive device renders it possible for the first time to measurean entire eye with the aid of the OCDR method in an A-scan with anaccuracy of better 100 μm, particularly better than 30 μm, in order thusto obtain measured values for matching an intraocular lens. Here, themeasurement can comprise two or more simultaneous distance measurementsbetween the cornea, lens and retina, and is robust against the usualaxial and lateral patient movements, which typically lie in the regionof 1 mm/s.

BRIEF DESCRIPTION OF THE DRAWINGS

The invention is explained in more detail below with the aid ofdrawings, in which

FIGS. 1, 1 a, 1 b and 1 c depicts a basic design of the inventivedevice,

FIGS. 2, 2 a, 2 b, 2 c and 2 d depicts an example embodiment of theinvention, and

FIGS. 3 a, 3 b, 3 c, 3 d and 3 e depicts various solutions for arrangingthe reference plane for the measurement.

DETAILED DESCRIPTION

The basic design for implementing the invention in FIG. 1 consists of asuitable tunable laser 1 that is characterized by the followingvariables: tuning time τ, wavelength λ, spectral tuning range Δk,centroid wave number k₀, and laser line width δk.

A beam shaping and coupling unit 2 serves both to direct the beam of thelaser 1 onto the sample 3 (illustrated schematically here as an eye),and to feed the light backscattered by the sample 3 to a detector 4, Din this case being the diameter of the measurement beam when it impingeson the sample (here the cornea of the eye). Assigned to the laser 1 is areference interferometer 5 whose preferred two detectors 6, 7 areconnected to a data acquisition apparatus 9 via a difference amplifier8. The detectors consist of InGaAs photodiodes with chip diameters >0.1mm and bandwidths less than 80 MHz.

Here, the reference interferometer 5 is a fiber optic Mach-Zehnderinterferometer having two outputs connected to the detectors 6 and 7.The optical path length difference between the arms of the referenceinterferometer is LREF. The arms of the reference interferometer 5 canconsist of fibers with various chromatic dispersions so that the samedispersion is implemented in both arms despite wavelength difference, inorder to attain a maximum modulation contrast at the detectors 6 and 7.It is also possible in principle to detect signals with the aid of onlyone detector 6 and without the difference amplifier 8, the signalquality being reduced.

The data acquisition apparatus 9 is likewise connected to the detector 4of the light backscattered by the sample 3, and to the control andevaluation unit 10. This connection permits the control and evaluationunit 10 to acquire and process the signals received from the detector 4and recorded with the aid of the data acquisition apparatus 9, and alsopermits synchronization of the measurement cycle between the dataacquisition apparatus 9 and control and evaluation unit 10. Furthermore,there is provided between the data acquisition apparatus 9 and laser 1 asynchronization connection 11 with the aid of which the tuning of thelaser 1 is synchronized with the data acquisition. The data acquisitionapparatus 9 is thereby also capable of transmitting control signals tothe laser 1 or of receiving them therefrom, if the laser carries outperiodic tuning independently.

As illustrated in FIG. 1 a, in this case the beam shaping and couplingunit 2 can be designed with a fiber coupler 12, while an implementationwith a beam splitter 13 is also possible, as shown in FIG. 1 b.

The control and evaluation unit 10 controls the tuning of the laser 1via the data acquisition apparatus 9 (with a spectral tuning range Δk inthe tuning time τ), and the light backscattered by the sample 3 andmeasured by the detector 4 is digitized and subjected in a known way toa Fourier transformation, for example a discrete Fourier transformationDFT, for the reconstruction of the A-scan.

In this case, the reconstruction of the A-scan is particularly performedin such a way that the interpolation point spacing in the A-scan issmaller than 4*ln 2/Δk, in particular smaller than 4/3*ln 2/Δk. If thecoupling element 2 does not include a further apparatus for generating areference light component, an autocorrelation signal is obtained as anA-scan, in a manner similar to that described by K. Hitzenberger, A. F.Fercher and M. Kulhavy in “Wavelength-tuning interferometry ofintraocular distances”, Appl. Optics 36 (1997), pages 6548 to 6553. ThisA-scan, which is symmetrical along the light propagation axis about thestrongest signal component (mostly the cornea surface signal) can betrimmed so that redundant information is omitted (FIG. 1 c).

The reference interferometer 5 serves in this case to determine thetiming of the tuning process exactly. To this end, a simultaneousrecording of the reference interferometer signal and measurement signalwith the aid of the data acquisition apparatus 9 permits theinstantaneous, relative wave numbers of the laser to be assigned veryexactly to the measurement signals.

In particular, it is possible thereby to interpolate measurement signalsfor any interpolation point systems in the spectral domain (remapping),for example in order, with reference to equidistant wave numbers intospecific tissues, to obtain A-scans with a constant resolution over theentire measurement depth range after the Fourier transformation.Furthermore, a numerical compensation of the chromatic dispersionreducing the resolution is also possible thereby. Such a remapping andthese dispersion compensations are described in U.S. Pat. No. 7,330,270for SD-OCT, the entire content thereof is incorporated by reference.

As described in V. J. Srininivasan et al. Optics Letters 32, 361,equidistant wave numbers can be obtained by way of example in this casevia a phase reconstruction of the signals from the referenceinterferometer 5 by means of Hilbert transformation with subsequentdetermination of equidistant phase spacings.

It is further necessary to this end to ensure that statistical errors inthe tuning of the wave number of the laser source 1 are so small thatthe phases of the signals from the reference interferometer 5 can beuniquely reconstructed. It is to hold true, particularly, that thestandard deviation of the wave number error σ_(k) remain below a limitσ_(k)<π/(LREF*v) prescribed by the length LREF, so that phase errors>πcan, for example, occur only every 400 (v=4.05) . . . 10 000 (v=3.23)scans.

The OCDR signals can also be improved by averaging and subtractingbackground signals that can be obtained without a sample, or with themeasurement beam masked out.

FIG. 1 c shows the A-scan obtained for the example of an eye as sample3, that is to say the longitudinal profile along the light axis. Thepeaks show from left to right in this case the reflections of thecornea, front side of the lens, rear side lens and retina.

FIG. 2 shows an example embodiment of the invention for biometry of theeye. The basic design consisting of a tunable laser source 1, beamshaping and coupling unit 2, reference interfometer 5 with the detectors6, 7, the assigned difference amplifier 8 and the data acquisition unit9 corresponds to the design known from FIG. 1. In addition, there isarranged downstream of the laser 1 a further interferometer 14, whichimplements interference of light shone into the sample 3 (eye) andbackscattered with a reference light component that can then be detectedwith the aid of the detectors 15 and 16. Assigned to the two detectors15, 16 of the interferometer 14 is an amplifier 17 that can, in aparticular embodiment, be switched over between difference amplificationand aggregate gain. The output of the amplifier 17 is connected, inturn, to the data acquisition unit 9.

The fundamental signal processing, that is to say, by way of example,the detection of the signals from the reference interferometer 5, theremapping, the background subtraction and the Fourier transformation, isperformed here in the same way as described previously for the obtainingof the auto correlation signals.

In this case, the reference light component is generated by means of atransmissive reference arm in an interferometer 14 consisting of fiberoptic components. Alternatively, it is also possible to use referencelight components obtained via a reflection, for example at fiber ends.FIGS. 2 a and 2 b show two variants of such a fiber interferometer. Anadjustable attenuator 26 that permits the reference light component tobe set for a favorable response of the detectors 15 and 16 is used,furthermore, in the reference arm of these interferometers. Possiblepolarization compensators in the sample arm or reference arm thatenables a partial or complete compensation of the polarization states ofthe reference light and sample light for the purpose of adequateinterference, are not illustrated.

The first fiber coupler 27 serves the purpose of division into a samplearm and reference arm.

The division ratio of the fiber coupler 27 is set in this case in FIG. 2a so that the predominant portion of the light returning from the sampleis fed to the detectors 15 and 16.

The fiber coupler 28 serves the purpose of interferometricallysuperposing the reference light from the reference arm on the returningsample light, and of feeding the signal components of opposite phase tothe detectors 15 and 16 for balanced detection. In principle, it is alsopossible to use the options, described in WO 2004/111 661, to determinefurther quadrature components for mirror artifact supposition inFD-OCDR.

In FIG. 2 b, the fiber coupler 27 effects a part of the reference armattenuation, and the fiber coupler 29 is selected in such a way thatlight returning from the sample is fed predominately to the detectors 15and 16. This is advantageous when the adjustable attenuator 26 has onlya limited attenuation range. Furthermore, visible light can still be fedin from the fixation light source 19 via the fiber coupler 27.

If the adjustable attenuator 26 is designed so that the reference lightcan also be completely blocked, it is consequently possible thereby alsoto obtain the above-mentioned autocorrelation signals. In this case, theamplifier 17 is switched over to summation while, given use of thereference light component for a balanced detection, it is necessary toswitch over to subtraction. Given a reference arm that is not blocked,summation signals can also be used for laser monitoring.

It has proved to be effective for the measurement of the eye to select ameasurement wavelength λ in the infrared region, that is to say, forexample, between 600 and 1150 nm, values such as 700 nm, 800 nm and 1060nm being particularly preferred.

Furthermore, the arrangement has an observation camera 18 and a fixationsource 19. The observation camera 18 corresponds to the CCD camera in WO00/33729, the entire contents of which is incorporated herein byreference, and serves, as there, to control the alignment of themeasurement arrangement relative to the eye 3 of the patient; to thisend, it is also possible in addition to project aiming markers onto theeye. This means that they should have at least a residual sensitivityfor the measuring wavelength λ and this is generally valid for cameraswith the silicon sensor (including in a CMOS fashion), in particular thecomparatively strong reflection of the measurement beam at the corneabeing effectively determinable. The fixation source 19 serves to alignthe eye relative to the measuring arrangement and can, for example, bedesigned in analogy with the proposal in DE 103 23 920, the entirecontents of which is incorporated by reference. The patient can beinfluenced by exerting appropriate control so that measurement isperformed optionally on the axis of the eye or the visual axis(“refixation”).

For the measurement, the patient places his head 20 onto the testsubject rest 21, in order for it to be held largely at rest. The laser 1and interferometer 14 are interconnected in a stable position and can bedisplaced jointly with the aid of means (not illustrated here) in orderto adjust with reference to the patient's eye 3. This stable positionconnection also proves to be advantageous when the light is guided viafibers, in particular monomode fibers. It is thereby particularlypossible to dispense with the use of variable fiber optic polarizationcompensators (paddles).

The data acquisition unit 9 permits the digitization of signals from theamplifier 8 (signal from the reference interferometer 5) and from theamplifier 17 (signal from the interferometer 14) with different bitdepths.

In order to reduce the data volume to be transmitted, the minimum bitdepth MBT for the digitization of the reference interferometer signal isin this case matched to the optical wavelength difference LREF in thereference interferometer 5 and to the optical OCDR measuring range ZMAXin the following way: MBT>=log 2(2/(1−|cos(π*LREF/2ZMAX)|). Given anLREF=100 mm and a ZMAX=60 mm, what is MBT=4 bit.

Bit depths of <14 bit are used to digitize the OCDR signals from theinterferometer 14, a bit depth of 10 bit, in particular, being adequate.

The control unit 10 is connected to the data acquisition unit 9 via aline 22 in order to enable a switchover of the bit depth of the measuredsignal. Such a bit depth switchover is, for example, advantageous for anoptimum response upon switchover between autocorrelation measurement andmeasurement with the aid of the reference arm.

Furthermore, the control unit 10 is connected to the beam shaping andcoupling unit 2 (line 23) in order to be able to implement a switchoverof the focus (as in DE 103 23 920), while the amplifier 17 can becontrolled via a further line 24, in particular can be switched overbetween the summation function and difference function. In addition, aswitchover of the amplification ratio of the amplifier is performed,particularly given overdriving of the OCDR signals, or in the event of aswitchover between autocorrelation measurement and measurement withreference to a reference light component from a reference arm.

For focusing purposes, the coupling unit 2 includes a focusing lens 30that generates a measurement beam focus in the eye posterior to thefront eye section (preferably 8 to 25 mm behind the cornea, at most 40mm behind the cornea). Alternatively, it is also possible to usediffractive or reflective focusing elements. An adaptation to therefraction of the eye can also be performed in this case by suitableadaptation optics. It is advantageous here when the fixation source 19is fed into the coupling unit 2 in such a way that the fixation is notinfluenced by the focusing of the measurement beam, and the bestpossible imaging of the fixing light on the retina results, that is tosay preferably downstream of the focusing lens 30.

Furthermore, the coupling optics can also include an apparatus fordisplacing or deflecting the measurement beam, in particular in order toimprove OCDR signals, for example by avoiding strong interfering cornealreflections by lateral measurement beam displacement.

The switching over of the attenuation or blocking of the reference armin the interferometer 14 is performed via a connection 25 between thecontrol unit 10 and interferometer 14. FIG. 2 c shows an A-scan that wasmeasured with the aid of an arrangement according to FIG. 2 (withreference arm), while FIG. 2 d shows an analogous measurement with ablocked reference arm, that is to say by means of determining theautocorrelation function of the backscattered sample light, withoutrepresentation of the mirror-symmetric signal component.

With the preferred values of tuning range Δk=112 000 m⁻¹, D=2 mm,wavelength λ=1060 nm and tuning time τ=500 μs, it is possible for thefirst time to determine the overall eye length and the position of thecrystalline lens in one measuring operation with an OCDRresolution/measuring accuracy of <30 μm. It is ensured here that themeasurement result is not corrupted by involuntary eye movements.

It has emerged that the maximum laser line width δk depends on themeasurement regime. The laser line width must be smaller than 162 m⁻¹ inthe case of measuring the autocorrelation function, that is to say witha blocked reference arm.

If use is made of an arrangement with a reference arm, it is possible bysuitable definition of the reference plane in the sample to ensurethat: 1. the signals from the cornea, lens and retina are detected withsufficient strength, and that 2. mirror artifacts can be suppressed oridentified and eliminated by computation. FIG. 3 shows thisschematically for various positions of the reference plane, which can beset via the wavelength differences between the reference and samplelight paths.

In FIG. 3 a, the reference plane was set behind the retina (R), theresult being a maximum laser line width of 93 m⁻¹ (signal drop of 80 dBover a total measuring range of 54 mm, represented schematically as acurve). R′C′ denote the mirror artifacts here and below.

In FIG. 3 b, the reference plane is set in front of the cornea (C), theresult being a maximum laser line width of 81 m⁻¹ (possible signal dropof 60 dB over a total measuring range of 54 mm, since the corneareflects better than the retina).

In FIG. 3 c, the reference plane was set between the cornea (C) andretina (R), the result being a maximum laser line width of 162 m⁻¹.

As in FIG. 3 a, in FIG. 3 d the reference plane was set behind theretina (R); given a target signal drop of only 20 dB over a totalmeasuring range of 54 mm, the result is a maximum laser line width of 47m⁻¹. This is the preferred laser line width δk. It is necessary here toimplement a minimum space of 64 mm between the reference planes and theoptical element that lies closest to the eye and is traversed by themeasurement beam.

FIG. 3 e shows the conditions in the case of the measurement of theautocorrelation function. For substantially narrower laser line widths(<20 m⁻¹), there is the risk that signal artifacts could occur fromreflections by the optical elements and corrupt the interpretation ofthe measurement results.

If the aim is to make joint use of measurement and mirror signals (FIG.3 c), or if there is a risk of undesired overlap between the measurementsignals (for example, R, C in FIG. 3 a, b), unique identification of themirror signals (R′, C′) is required to avoid error.

It is preferred for this purpose to implement variations in theconditions of measurement or reconstruction in such a way as to performunique variations in the mirror signal that permit a manual andautomated identification, or else suppression.

A first example option for this is the variation in the differencebetween the lengths of reference arm and sample arm, which in the caseof the mirror signal effects a local variation that is opposite to themeasurement signal and can be detected numerically. This can beperformed, for example, by varying the distance between the test subjectand measuring apparatus, preferably in the range of 0.1 to 4 mm, and bydetermining the first moment of the signal distribution of the signalcomponent to be identified. Alternatively, it is also possible toperform between successive tunings of the laser very quick phase shiftsbetween the reference arm and sample arm by means of which it ispossible to use the spectral data to reconstruct a complex FD-OCDRsignal in the case of which it is also possible to suppress the mirrorartifact completely or partially (compared to U.S. Pat. No. 7,433,046B2). This complete or partial suppression of mirror artifacts is thenalso suitable as an identification feature. However, this procedure ismore complicated and more sensitive to sample movements than thedetermination of relatively large local variations. A further option isthe use of an unbalanced chromatic dispersion in accordance with U.S.Pat. No. 7,330,270, there subsequently being, however, a need to performnumerical reconstruction several times with different dispersioncompensation coefficients in order to effect and determine differentdegrees of smearing of the measurement and mirror signals. Inparticular, a change in sign during the dispersion correction has theeffect that smearing takes place from the mirror signal to themeasurement signal, as a result of which it is likewise possible toperform an identification, for example by determining the second momentof the distribution of the signal component to be identified.

A further advantageous solution is to fundamentally avoid an undesiredoverlap between the measurement and mirror signals in such a way as toexclude a positioning of the interferometric reference plane in theregion between the cornea and retina. One possible embodiment to thisend is to design the interferometer so that the reference plane isalways reliably positioned in front of the cornea of a test subject heldin a head support (FIG. 3 b), while the measuring range is selected tobe sufficiently large in order to cover the positioning error and theeye length range, particularly by selecting an interferometric measuringrange (one-sided A-scan length) of at least 44 mm. If the aim is tofacilitate adjustment of the patient's eye, it is recommended to selectan interferometric measuring range of at least 50 mm.

It is advantageous that in the case of the inventive solution it ispossible not only to detect a single eye region, such as the frontchamber or retina, by a B-scan, but that it is also possiblesimultaneously to obtain additional B-scan information from therespective other regions. This is particularly advantageous, since thisadditional information permits not only very accurate measurements ofthe spacing of the eye structures inside individual A-scans despitepossible eye movements over the duration of the B-scan recording, butalso allow the correction of the B-scan data for a reproduction ofmovement corrected image data. By way of example, it would be preferredfor this purpose to record a B-scan of a retina region while the B-scansimultaneously includes image information of the cornea such as, forexample, specular reflections from the front or rear surface of thecornea. Were there allowed to be axial eye movements of the test subjectduring the B-scan recording, these movements would jointly influence theprofiles of the corneal and retinal structures. If the image informationof the cornea is now corrected in such a way that it would correspond toa shape otherwise determined by measurement or recorded, the correctionused to this end can also be applied to the image informationcorresponding to the retinal structures, and a movement correctedrepresentation can likewise be attained thereby. By way of example, sucha correction can be implemented by axial and lateral displacements ofthe A-scan forming the B-scan, particularly in such a way that thecorneal structures in the image exhibit, for example, minimum deviationsfrom a shape that is determined in advance by corneal keratometry orcorneal topography and is, in particular, to be described by continuousfunctions.

An alternative correction option is to analyze the cornea profile withreference to its apparent local surface curvatures in the B-scan, and toderive modulation frequencies that can then be separated into componentsof shape and movement by plausibility considerations in ordersubsequently to reduce the movement component in the image data. Thiscan preferably take place in such a way that at least one boundarysurface of the cornea is determined, and the axial positions thereof areplotted relative to the lateral deflection or else the correspondingA-scan recording instances, and are transformed into a modulationfrequency spectrum by Fourier transformation. Subsequently, shape andmovement components can be obtained by a suitably fashioned filter. Themovement components can subsequently be corrected by Fourier backtransformation and oppositely directed displacement of the A-scans inthe image data so that the cornea and retina as well as the frontchamber appear with corrected movement. These abovementioned filters forthe modulation frequencies can be obtained, for example, by analysis ofclinical topography data, in a way known per se from U.S. Pat. No.7,370,969, the entire content of which is hereby incorporated byreference.

The implementation of the invention is not tied to the exampleembodiment illustrated—expert developments do not go beyond the scope ofprotection of the patent claims.

The invention claimed is:
 1. A device for swept source optical coherencedomain reflectometry (SS OCDR) on a movable sample, for the purpose ofobtaining A-scans, comprising: a tunable laser light source; at leastone receiver for light backscattered from the sample, wherein the sampleis illuminated on the sample surface with a measurement beam of diameterD; wherein the light source has a line width of δk<168 m⁻¹, wherein theA scan measuring range corresponds to a depth of the sample and furtherwherein the device is configured such that tuning of the light source isperformed in τ<44 sec/(D*k₀) about a centroid wave number k₀.
 2. Thedevice for SS OCDR as claimed in claim 1, wherein the sample comprises ahuman eye.
 3. The device for SS OCDR as claimed in claim 2, wherein themeasurement beam is convergent before entering the eye.
 4. The devicefor SS OCDR as claimed in claim 3, wherein a degree of the convergenceis tuneable or switchable.
 5. The device for SS OCDR as claimed in claim2, wherein the measurement beam is colliminated before entering the eye,and further comprising means for refixation of the eye so thatmeasurement is made on one of an optical axis of the eye and a visualaxis of the eye.
 6. The device for SS OCDR as claimed in claim 2,wherein the measurement beam is switchable between a colliminated and aconvergent measurement beam.
 7. The device for SS OCDR as claimed inclaim 2, wherein the measurement beam strikes the eye outside a cornealapex.
 8. The device for SS OCDR as claimed in claim 7, furthercomprising an apparatus that controls positioning the measurement beamrelative to the eye.
 9. The device for SS OCDR as claimed in claim 8,wherein the apparatus that controls the positioning of the measurementbeam relative to the eye does so by evaluating the light detected by thereceiver.
 10. The device for SS OCDR as claimed in claim 2, wherein aphoton flux of 3*10⁸ to 3*10¹³ is directed onto the eye during tuning oftime τ of the light source.
 11. The device for SS OCDR as claimed inclaim 2, wherein measurement is performed along a visual axis of theeye.
 12. The device for SS OCDR as claimed in claim 2, wherein detectionof light backscattered by a cornea, a crystalline lens and a retina isperformed during a tuning of the light source.
 13. The device for SSOCDR as claimed in claim 2, wherein reference signals and sample signalsare digitized at constant scanning rates.
 14. The device for SS OCDR asclaimed in claim 13, wherein the constant scanning rates for referencesignals and sample signals are the same.
 15. The device for SS OCDR asclaimed in claim 2, wherein the light source is movable relative to theeye.
 16. The device for SS OCDR as claimed in claim 15, furthercomprising a reference interferometer and wherein the light source andthe reference interferometer are rigidly connected.
 17. The device forSS OCDR as claimed in claim 16, wherein the reference interferometer hasa wave number reference for length calibration of the OCDR signal. 18.The device for SS OCDR as claimed in claim 2, further comprising aninterferometer that is movable relative to the eye, the light source andinterferometer being rigidly connected.
 19. A device for SS OCDR asclaimed in claim 2 further comprising means for variation in one ofmeasurement conditions and reconstruction conditions, the variation inone of the measurement conditions and the reconstruction conditionsbeing undertaken so as to attain one of identification and suppressionof a mirror artifact signal.
 20. The device for SS OCDR as claimed inclaim 19, wherein the means for variations in one of measurementconditions and reconstruction conditions comprises a unit for varying apath length difference between the sample arm and reference arm of theinterferometer.
 21. The device for SS OCDR as claimed in claim 1,wherein the light source has a spectral tuning range Δk about a centroidwave number k₀ of at least Δk>18 000 m⁻¹.
 22. The device for SS OCDR asclaimed in claim 21, wherein the ratio of a tuning range Δk and linewidth δk is greater than
 360. 23. The device for SS OCDR as claimed inclaim 21, wherein a quotient of the tuning rate Δk/τ and laser linewidth δk is greater than 18 kHz.
 24. The device for SS OCDR as claimedin claim 21, wherein the backscattered light detected at the receiver isdigitized at a rate of more than Δk/(τ*δk).
 25. The device for SS OCDRas claimed in claim 1, wherein the line width δk of the light sourcelies between 22 and 50 m⁻¹.
 26. The device for SS OCDR as claimed inclaim 1, wherein the bandwidth of the at least one receiver is greaterthan 2*Δk/(τ*δk).
 27. The device for SS OCDR as claimed in claim 26,wherein the bandwidth of the at least one receiver is less than 80 Mhz.28. The device for SS OCDR as claimed in claim 1, further comprising aninterferometer with a reference arm.
 29. The device for SS OCDR asclaimed in claim 28, wherein the position of retina signals and corneasignals in the A-scan and the laser line width 6 k are matched to oneanother.
 30. The device for SS OCDR as claimed in claim 1, wherein themeasurement beam diameter D is smaller than 3 mm.
 31. The device for SSOCDR as claimed in claim 1, further comprising monomode fibers in atleast one of a reference arm, a source arm, a detection arm, and areference interferometer.
 32. The device for SS OCDR as claimed in claim1, wherein the measurement beam has a wavelength of between 600 and 1150nm.
 33. The device for SS OCDR as claimed in claim 1, further comprisingan apparatus for projecting onto the eye aiming markers with awavelength of between 400 and 900 nm.
 34. The device for SS OCDR asclaimed in claim 33, further comprising a camera for checking adjustmentof the measurement beam relative to the eye.
 35. The device for SS OCDRas claimed in claim 34, wherein the camera is sensitive to thewavelengths of the measurement beam and the aiming markers.
 36. A methodfor SS OCDR on the eye, comprising: detecting and evaluating lightbackscattered by a cornea from individual tunings of a light source withregard in each case to time dependent cornea positions; comparing thetime dependent cornea positions with an expected profile; and usingdeviations from the expected profile to correct position data of cornea,crystalline lens and/or retina.
 37. A method for SS OCDR on the eye, asclaimed in claim 3, further comprising deflecting a measurement beamlaterally at one of during individual tunings and between the individualtunings.
 38. A method for SS OCDR on the eye, as claimed in claim 36,further comprising illuminating the eye with a measurement beam ofdiameter D from a tunable laser light source; and detecting lightbackscattered by a cornea, a crystalline lens and a retina during atuning of the tunable laser light source wherein the light source has aline width of δk<168 m⁻¹, and further wherein tuning of the light sourceis performed in τ<44 sec/(D*k₀) about a centroid wave number k₀.